When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the nuclei in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) that is in the x-y plane and that is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped,” into the x-y plane to produce a net transverse magnetic moment Mxy. A signal is emitted by the excited nuclei or “spins,” after the excitation signal B1 is terminated, and this signal may be received and processed to form an image.
When utilizing these “MR” (magnetic resonance) signals to produce images, magnetic field gradients (Gx, Gy, and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received MR signals are digitized and processed to reconstruct the image using one of many well-known reconstruction techniques.
The measurement cycle used to acquire each MR signal is performed under the direction of a pulse sequence produced by a pulse sequencer. Clinically available magnetic resonance imaging (MRI) systems store a library of such pulse sequences that can be prescribed to meet the needs of many different clinical applications. Research MRI systems include a library of clinically-proven pulse sequences and they also enable the development of new pulse sequences.
The MR signals acquired with an MRI system are signal samples of the subject of the examination in Fourier space, or what is often referred to in the art as “k-space.” Each MR measurement cycle, or pulse sequence, typically samples a portion of k-space along a sampling trajectory characteristic of that pulse sequence. Most pulse sequences sample k-space in a raster scan-like pattern sometimes referred to as a “spin-warp,” a “Fourier,” a “rectilinear,” or a “Cartesian” scan. The spin-warp scan technique employs a variable amplitude phase encoding magnetic field gradient pulse prior to the acquisition of MR spin-echo signals to phase encode spatial information in the direction of this gradient. In a two-dimensional implementation (“2DFT”), for example, spatial information is encoded in one direction by applying a phase encoding gradient, Gy, along that direction, and then a spin-echo signal is acquired in the presence of a readout magnetic field gradient, Gx, in a direction orthogonal to the phase encoding direction. The readout gradient present during the spin-echo acquisition encodes spatial information in the orthogonal direction. In a typical 2DFT pulse sequence, the magnitude of the phase encoding gradient pulse, Gy, is incremented, ΔGy, in the sequence of measurement cycles, or “views” that are acquired during the scan to produce a set of k-space MR data from which an entire image can be reconstructed.
MRI can be used to measure the exchange of magnetization between molecules to provide unique information about the chemical and molecular environment of samples or tissues. One type of such exchange measurement is broadly referred to in the field as magnetization transfer. This technique is capable of measuring the exchange of magnetization from spin species that have short transverse relaxation times (T2). Because many different molecules have short T2, this technique is not particularly sensitive to specific molecules.
A second type of magnetization exchange occurs between water protons and a molecule with long enough T2 that its difference in frequency from water can be observed. Saturation of the magnetization from this molecule will generally decrease the measurable signal from water. This effect is generally referred to in the field as chemical exchange saturation transfer (“CEST”). Two different types of molecules can generate CEST effects: endogenous, or naturally occurring, molecules and exogenous contrast agents. In either instance, the molecules whose chemical exchange with water produces the CEST effect are generally referred to as so-called “exchangeable protons.”
The CEST imaging method offers three advantages over traditional molecular MRI techniques. First, in some cases the molecules of interest within the subject can be directly detected. This feature mitigates the need for administering contrast agents to the subject. Second, the image contrast mechanism can be controlled with the RF pulses produced by the MRI system and, as such, can be turned on and off when desired. This control allows the location of specific molecules of interest to be detected by comparing images having the desired contrast present to those where it has been turned off. Lastly, the CEST imaging method is far more sensitive than traditional molecular MRI techniques, making it able to detect substantially low concentrations of given molecules.
Thus, CEST MRI is a sensitive imaging technique for detecting compounds containing exchangeable protons. Such labile protons can be selectively saturated by an RF pulse, and the saturation subsequently transferred to the bulk water signal via proton chemical exchange, resulting in substantial sensitivity enhancement. CEST imaging has been demonstrated in mapping low-concentration metabolites such as creatine (Cr), glucose, glutamate, and changes in microenvironment properties such as temperature and pH, promising a host of in vivo applications such as imaging of ischemic stroke and tumor. One example is amide proton transfer (APT), a specific form of CEST imaging that is sensitive to pH and mobile protein/peptide content. Recently, APT imaging has shown promising results in characterizing tissue acidosis after ischemic stroke, tumor detection, grading, and differentiation of tumor recurrence from radiation necrosis.
Despite holding promises in stroke and tumor imaging, CEST MRI suffers from its qualitative nature, which depends on many factors, including the chemical exchange rate, concentration of exchangeable protons, longitudinal relaxation time, and RF saturation power. To quantify the CEST effect relative to, for example APT, an asymmetry analysis (MTRasym) is most commonly used to suppress interference from non-linear direct water saturation (RF spillover) and broadband magnetization transfer (MT) effects by taking the difference between a reference image (e.g. −3.5 ppm) and label imagine (e.g. +3.5 ppm). Note the term ‘direct water saturation’ is used interchangeably with ‘RF spillover’ or ‘direct saturation’ or ‘spillover’ in this context. However, in vivo MTRasym is contaminated by asymmetric magnetization transfer contrasts (MTC) from semisolid macromolecules and nuclear overhauser enhancement (NOE) effects, which result in a negative and inhomogeneous shift across the brain. As such, in vivo MTRasym has a mixed contribution from APT, asymmetric NOE and MTC effects. If the asymmetric background signals are not removed the result is a substantial underestimation of the CEST effect. Another drawback associated with the MTRasym method is that MTRasym analysis is subjected to static magnetic (B0) field inhomogeneity, which needs additional acquisition of densely sampled Z-spectrum to correct for B0 inhomogeneity using interpolation approach. The dominant B0 inhomogeneity effect also stems from the non-linear direct saturation effect.
Therefore, it would be desirable to develop quantitative methods that can determine in vivo CEST effects without the need of time-consuming Z-spectrum acquisition while controlling for confounding contaminations.